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16th June 2015
Orthotic Heel Wedges do not alter Hindfoot Kinematics and Achilles Tendon Force
during Level and Inclined Walking in Healthy Individuals
Robert A Weinert-Aplin,1,2 Anthony M J Bull,2 Alison H McGregor1
1 Department of Surgery and Cancer, Imperial College London, London, U.K.; 2 Department
of Bioengineering, Imperial College London, London, U.K.
Funding: This study was funded by an EPSRC Case award, with financial contributions from
Vicon Motion Systems for the running costs of the project.
Conflicts of interest statement: None
Correspondence Address:
Robert Weinert-Aplin, Department of Bioengineering, Royal School of Mines, South Kensington Campus, Imperial College London, London, SW7 2AZ, United Kingdom
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Abstract
Conservative treatments such as in-shoe orthotic heel wedges to treat musculoskeletal
pathologies are not new. However, the mechanical basis by which such orthoses act have not
been elucidated. Quantifying the mechanical changes that occur when wearing heel wedges
may help to explain the mixed evidence supporting their use in management of Achilles
tendonitis.
A musculoskeletal modelling approach was used to quantify changes in lower limb
mechanics when walking due to the introduction of 12mm orthotic heel wedges. A control
group of 19 healthy volunteers walked on a level and inclined walkway while optical motion,
forceplate and plantar pressure data were recorded as model inputs.
Heel wedges induced a posterior shift of centre of pressure that resulted in increased ankle
dorsi-flexion moments and reduced plantar-flexion moments. Consequently, this resulted in
increased peak ankle dorsi-flexor muscle forces during early stance and reduced Tibialis
Posterior and toe flexor muscles forces during late stance. Heel wedges did not reduce triceps
surae muscle forces during any walking condition.
These results add to the body of clinical evidence that does notagainst the use of support heel
wedges hypothesised to reduce Achilles tendon loading, as a means to treat Achilles
tendonitis according to the hypothesis that heel wedges reduce Achilles tendon load during
walking . our and the findings provide an explanation as to why this may be the casetheory is
not appropriate.
Keywords: tendonitis, tendinitis, musculoskeletal, modelling, conservative treatment
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Introduction
The Achilles tendon is functionally important as it is the main driver of ankle motion
during locomotion. Consequently it is a highly loaded tendon, with reported loads of up to 5
times body weight (BW) during level walking1 and 8.2BW- 12.5BW during running.2,3 With
such high cyclic loads being experienced by the Achilles, it is unsurprising that the Achilles
is a common site of overuse injury,4 with 5-18% of all lower extremity injuries involving the
Achilles tendon.5 As a result of this high injury prevalence, there have been a number of
reviews into Achilles injuries.6-11
Risk factors for Achilles injuries include: magnitude of Achilles tendon load,
inappropriate equipment such as inflexible shoes, low heel tabs on running
shoesinappropriate footwear12, training errors7 and abnormal kinematics.6 Abnormal
kinematics are generally considered to be related to over-pronation of the subtalar joint
hindfoot causing asymmetric loading across the Achilles tendon.7 While there is some
evidence to show differences in strain across the tendon cross-section13 and along the length,14
over-pronation alone has not been linked to injury risk in running.5,15-17 It has also been shown
that differences exist between Medial Gastrocnemius and Soleus contractile behaviour during
walking,18,19 running20 and cycling,21 suggesting that differential strains across the tendon may
exist naturally during these activities.
Treatments for tendinopathies have varied substantially over the years, but have
always been aimed at reducing or eliminating pain in the tendon. Conservative treatments
which directly address the Achilles tendon include: have included: Rest Ice Compression
Elevation, eccentric strengthening exercises,22-25 insoles and splints24,26 and ultrasound.27 Of
the treatments that influence mechanical alignment of the hindfoot, insoles and splints are the
only option24,26, but have been shown
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The use of insoles or orthotics to correct abnormal hindfoot motion and correct for
biomechanical mal-alignments to reduce pain and aid in return to sport has been shown to be
possiblefollowing Achilles tendinopathy.17 However, the direct link between correcting
hindfoot motion and tendon healing has not been established, but this treatment is still
recommended.9
Currently it is believed that incorporating heel wedges aids in reducing tendon strain
during activities to avoid excess tendon loading and reduce pain during running.26,28,29
Investigations regarding heel wedges which plantarflex the ankle by raising the heel equally
on the medial and lateral sides, insoles and other orthotics have had variable success
regarding reduction of pain and ability to return to sport.6,20,26,28,30-33 However, as orthotics
require no invasive procedures and injury management can be at home and on-going, they are
particularly appealing. ThisA similar approach has been used in osteoarthritis, where knee
adduction moment has been the target of a variety of orthotics that aimed to alter the
mechanical loading of the knee by altering the frontal plane alignment of the hindfoot, with
several studies showing positive effects of using foot orthotics on knee adduction moment34
and tibio-femoral load.35 However, while studies investigating such orthotics have been
focussed on influencing out of plane moments at the knee, their investigations of orthotics
focussed on relieving Achilles tendon strain through plantarflexion of the ankle effect on the
ankle is often not reportedare less common. With wedges being able to alter loading at the
knee, it is not unreasonable to hypothesise that a similar approach could alter the loading at
the ankle, and indeed, this is the basis by which heel wedges are thought to operate during
walking when managing Achilles tendon pain in Achilles Tendonitis patients.36 Therefore, the
hypothesis of this study was that heel wedges which raise the heel are able to reduce Achilles
tendon force during level and inclined walking.
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The aim of this study was to quantify the effect of heel wedges on lower limb
mechanics, specifically ankle joint angles, moments and muscle forces; and relate these to
common injuries such as Achilles tendonitis. This is conducted during level and inclined
walking in order to understand their effect beyond the constrained context of level walking.
Methods
The subject participant group consisted of nineteen healthy individuals, with no
history of ankle injuries and no lower limb injury in the last 12 months and no clinical
symptoms of Achilles Tendinopathies (8 male [mean (SD); age: 28 (3); height: 1.76 m (0.10);
mass: 73.4 kg (12.0)] and 11 females [age: 29 (6); height: 1.63 m (0.05); mass: 58.7 kg
(10.2)]. Individuals were excluded if they had ever been diagnosed with Achilles
Tendinopathy or had any previous musculoskeletal or neuromuscular condition of the lower
limb. Ethical approval was obtained by a university institutional review board in accordance
with the Declaration of Helsinki and all participants were given an information sheet and
provided a signed consent form upon arrival.
A pair of commercially available orthotic heel wedges were used by all subjects
participants (Elevator Proheel™, Talar Made Orthotics Ltd, Springwood House, Foxwood
Way, Chesterfield, Derbyshire, England) (Figure 1). The wedges are made from medium
density ethylene vinyl acetate (EVA) foam and are designed to mould to the rearfoot and
elevate the heel, with the aim of being used as a tendonitis treatment. Wedges were available
for UK shoe sizes 2-5; 6-9 and 9.5-12.5 and all wedges had a 12mm rise from the front edge
of the wedge to approximately where the centre of the heel would be. Subjects Participants
were fitted with wedges corresponding to their running shoe size. All participants wore
standard running shoes that they felt comfortable in. Shoes were checked visually for
excessive wear under the sole and participants confirmed that they had used their running
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shoes previously before participating in the study. A wedge height of 12mm was chosen as it
represents a height that is recommended for Achilles tendonitis patients.36
The mechanical effect of heel wedges on a variety of walking conditions necessitated
the need for an inclined walkway which could securely accommodate a forceplate either on a
level or inclined surface, shown in (Figure 2). The setup allowed the subjects participants to
approach the 10° incline on a level surface for several steps, before ascending the 2m inclined
section where foot strike was recorded followed by a few steps of level walking at the top of
the incline. Subjects Participants were then asked to turn around and walk down the incline to
provide the data for the downhill walking condition. For both inclined and declined walking,
all subjects participants required three steps to cover the 2m inclined section, with the middle
step being used for subsequent analysis. Participants were given as long as they needed to
familiarise themselves with each condition. This was assessed by participants themselves as
they walked freely around the laboratory and up and down the level and inclined walkway
untiltill they felt comfortable. The amount of time taken by each individual was not
measured, but was on the order of a few minutes.
Subjects Participants were given time to familiarise themselves with the equipment
and testing protocol before data collection began. 3D Ooptical motion (Vicon Motion
Systems, Oxford, UK), plantar pressure (Novel GmbH, Munich, Germany) and forceplate
(Kistler, Winterthur, Switzerland) data were collected for all conditions and used as inputs to
the musculoskeletal model (described below). Optical motion and plantar pressure data were
recorded at 100Hz and forceplate data was were recorded at 1000Hz. Data were recorded
continuously while the subjects participants walked over the level or inclined walkway with a
minimum of 5 clean strikes of the forceplate when walking in running shoes (“shod
walking”) or in running shoes with the orthotic heel wedges (“wedged walking”). Both the
walking condition order (shod or wedged walking) and walking incline (level, uphill or
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downhill) were randomised. Subjects Participants were given time to become used to the heel
wedges when going between shod and wedged walking conditions on each incline and were
able to walk freely along the level or inclined walkway without targeting the forceplate.
The musculoskeletal model used here has been described elsewhere previously;37 in
summary, it is a unilateral model of the lower limb, scaled to subject participant height and
weight by optical markers placed on the pelvis and lower limb (Figure 3), with the Achilles
insertion, the first metatarsal head and base, the fifth metatarsal head and base and the tip of
the second phalanx digitised as virtual landmarks relative to the marker clusters on the foot
and hallux. Error: Reference source not foundThis Error: Reference source not foundThe
measured optical motion data was were used to calculate joint angles and inter-segmental
moments using Euler angle decompositions and Newton-Euler equations at the
metatarsophalangeal (MTP), ankle, knee and hip joints respectively and a static optimisation
routine used to estimate and muscle forces for the 13 muscles crossing the ankle joint and
Achilles tendon force is taken as the sum of the triceps surae muscle forces. The knee and hip
were modelled with 3 rotational degrees of freedom and the ankle modelled as a saddle with
2 degrees of freedom and the MTP as a hinge with 1 rotational degree of freedom. Inter-
segmental moment data are presented in the local segment coordinate frame in which they
were calculated. Centre of pressure (CoP) data is are presented as dimensionless values
normalised to foot length, defined as the RMS distance from the calcaneus to the second
metatarsal head.
Statistical analyses were performed in Matlab using the Statistical Analysis Toolbox
(Version 2010b, The Mathworks Inc.). All data was checked for normality using a
Kolmogorov-Smirnov test. As the effect of orthotic heel wedges on lower limb mechanics
was the parameter of interest, statistical comparisons between shod and wedged walking for
each incline individually was performed using paired t-tests, with the level of significance set
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at 0.05. P-values under 0.1 are presented as trend changes and values above 0.1 are not
presented. Statistical comparisons for all hip, knee, ankle and MTP kinematics and inter-
segmental moments were performed at heel-strike (HS), defined by a forceplate force
exceeding 40N, weight-acceptance (WA), push-off (PO) and toe-off (TO), with the latter
three time-points defined according to changes in knee flexion angle.38 Inter-segmental
moments were normalised to body weight (BW) and height (ht).39 Peak muscle forces were
compared across all of stance phase, and in the case of the inv/evertor muscles, peak forces
were compared during early (<50 %) and late (≥ 50 %) stance.
Results
For the full gait analysis results, the reader is directed to the supplementary material.
The results presented here are those directly related to the calculation of the ankle muscle
forces. Changes in overall gait dynamics due to heel wedges were only observed during
inclined walking as an increase in stance time (Table 1). During inclined walking, delays of
1-2 % stance to the first ground reaction force (GRF) peak when walking with heel wedges
was observed (P = .014 and P = .015 for uphill and downhill respectively). During uphill
walking an increase in stance time of 20ms was observed (P = .025) along with delays of 1 –
2 % stance to both GRF peaks (P = .014 and P = .028 respectively). Also reductions in the
magnitude of the first GRF peak (1 % BW, P = .027) and rate of force development (ROFD)
(6% reduction, P = .026) were observed during wedged uphill walking. Compared to shod
walking, the most anterior centre of pressure (CoP) position was found to be less anterior
during uphill and downhill wedged walking by 3% (P = .002) and 4% (P = .014) foot length
respectively.
At all time points considered no changes in peak frontal plane kinematics were
observed during level or inclined wedged walking, but changes in sagittal plane ankle
kinematics were observed during inclined walking (Figure 4). Flexion-extension angles at the
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ankle and MTP joints were largely unaffected due to heel wedges during level walking, with
only the MTP range of motion (ROM) decreasing significantly (17.7±5.4° vs. 15.4±3.5°, P =
.048). During inclined walking, the only cChanges in lower limb kinematicsankle angle were
confined to the ankle during early stance, with a more plantar-flexed ankle at HS (-1.4±5.5°
vs. -4.5±7.2°, P = .024 and -9.0±4.3° vs. -10.9±5.3°, P = .015 uphill and downhill
respectively) and WA (-1.0±5.5° vs. -5.2±9.4°, P = .034, -18.6±4.8° vs. -21.4±5.3°, P = .003
uphill and downhill respectively) during wedged walking. Sagittal plane ankle angles were
unaffected by heel wedges during level walking. During downhill walking, there was greater
knee flexion at WA (-28.1±5.8° vs. 30.0±5.5°, P = .020). Compared to the shod condition,
ankle ROM was observed to decrease during uphill wedged walking (35.1±5.7° vs.
32.1±4.6°, P = .035), but increase during downhill wedged walking (21.9±5.4° vs. 25.9±6.3°,
P < .001).
Knee, aAnkle and MTP joint moments were significantly affected by the presence of
heel wedges at each incline (Figure 5). Changes at the ankle during level walking were less
apparent compared to uphill or downhill walking, with delays to peak ankle dorsi-flexion and
plantar-flexion moments by 2% stance and 1% stance respectively only. However, a large
increase in knee push-off moment was observed (0.10 vs. -0.00 N m∙BW -1∙ht-1, P = .016)
during level wedged walking.
During uphill walking, peak ankle dorsi-flexion moment increased (0.01 vs. 0.06 N
m∙BW-1∙ht-1, P = .005), peak plantar flexion moment decreased (-0.85 vs. -0.79 N m∙BW-1∙ht-1,
P = .002) and peak inversion moment decreased (0.09 vs. 0.06 N m∙BW -1∙ht-1, P < .001). A
decrease in MTP flexion moment was also observed (-0.10 vs. -0.09 N m∙BW-1∙ht-1, P < .001).
At the knee, the first adduction peak moment was increased during wedged walking (0.16 vs.
0.20 N m∙BW-1∙ht-1, P < .001), while knee extensor moment at PO decreased (0.16 vs. 0.09 N
m∙BW-1∙ht-1, P < .001).
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Similar changes in ankle moments were observed during downhill walking, with an
increase in dorsiflexion moment (0.09 vs 0.14 N m∙BW-1∙ht-1, P < .001), a decrease in plantar-
flexion moment (-0.62 vs. -0.57 N m∙BW-1∙ht-1, P = .002) and a decrease in inversion moment
(0.09 vs. 0.07 N m∙BW-1∙ht-1, P = .030). Knee extensor moment was increased at WA (-0.64
vs. -0.71 N m∙BW-1∙ht-1), PO (-0.27 vs. -0.33 N m∙BW-1∙ht-1) and at TO (-0.39 vs. -0.47 N
m∙BW-1∙ht-1) during wedged walking (P < .001 for all three time-points).
The most consistent change in muscle force estimates due to heel wedges were in the
ankle dorsi-flexors (12–26 % BW increases in peak Tibialis Anterior force) and toe extensors
(range of 4–6 % BW increase for Extensor Digitorum/Hallucis Longus forces) muscle forces
during the first half of stance across all walking conditions (Figure 6 and Table 2). A second
consistent observation was the decrease in peak Tibialis Posterior and toe flexor forces during
level and inclined wedged walking, although this was only statistically significant during
inclined walking (mean decreases of 12-14 % BW for Tibialis Posterior and 9–11 % BW the
toe flexor muscles respectively). Critically, there was no statistically significant reduction in
peak Achilles force for any walking incline due to heel wedges (range of 5–14 % BW
decrease). The only significant changes in triceps surae loading during uphill walking were
decreases in the medial parts of the triceps surae (12 % BW and 6 % BW for medial Soleus
and medial Gastrocnemius respectively). During downhill walking, only the medial portion
of Soleus showed a significant decrease in peak force (9 % BW). Overall ankle joint reaction
force was reduced by 29 % BW during downhill wedged walking.
Discussion
The aim of this study was to quantify the effect of heel wedges on lower limb
mechanics, specifically ankle joint angles, moments and muscle forces; and relate these to
common injuries such as Achilles tendonitis. The main clinical driver behind assessing the
effect of orthotics on ankle loading during inclined walking was to determine how effective
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heel wedges are at reducing tendon load not only on during level walking, but also on
inclined surfaces, where Achilles tendon loads are known to be increased. Given the mixed
evidence surrounding the use of heel wedges to manage Achilles tendonitis, a broader
characterisation of lower limb mechanics due to orthotic heel wedges would provide some
insight into how the body may adapt to walking with such an intervention, allowing for
improvements regarding injury management.
A number of kinematic and kinetic changes were consistently observed across
walking conditions and these will be discussed further. However, the key finding of the study
was that 12mm orthotic heel wedges did not result in a reduction in peak triceps surae or
overall Achilles tendon forces, as was expected from current theories on the mechanism of
treating Achilles tendonitis with heel wedges.
No changes in ground reaction forces or ankle inversion kinematics were observed for
any wedged walking condition. While inferring subtalar kinematics through angles derived
from markers on the shoe has its limitations, information regarding out of plane changes in
kinematics can still be gained. However, as the cohort tested here consisted of healthy
individuals, hindfoot motion may not have required any correction, and as such the lack of
difference in hindfoot kinematics here should not automatically be extrapolated to a patient
population. This has similar implications for the kinetics, as kinetic changes are in response
to the kinematic and CoP changes that occur. If patient groups reported similar CoP and
kinematics changes, then it is likely that the kinetic results obtained here will have direct
relevance to the mechanical response of the lower limb to orthotic wedges in a patient
population. It should be noted that the observations here are valid only for this type of heel
wedge and wedges aiming to shift CoP in the frontal plane may act differently.
Changes in CoP were consistently shifted posteriorly across walking conditions. The
more posterior CoP at push-off coupled with relatively unaltered knee kinematics resulted in
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greater knee flexion moments (Figure 5), between weight-acceptance and push-off. This
increased demand on the knee joint likely has clinical implications for the quadriceps
muscles, as they are the main extensors of the knee, both in terms of fatigue of the muscles as
well as possible changes in knee joint loading. The posteriorly shifted CoP coupled This had
two distinct effects on lower limb loading. During early stance, a posterior shift in CoP would
place the GRF more posterior to the ankle joint centre, resulting in a greater dorsiflexion
moment and dorsiflexor muscle forces. During late stance at push-off, a posteriorly shift in
CoP places the GRF closer to the ankle joint centre, resulting in smaller plantar flexion
moments.with greater ankle plantarflexion from heel-strike until push-off would also have
resulted in greater dorsiflexion moments during weight-acceptance and smaller plantarflexion
moments at push-off. However, as was previously mentioned, this did not result in a
reduction in peak triceps surae or overall Achilles tendon forces. Instead, a redistribution of
the triceps surae loads was observed (Figure 6), where the medial triceps surae muscle loads
reduced and lateral triceps surae loading increased, while the peak forces of the less efficient,
secondary ankle plantar -flexor muscles of Tibialis Posterior and the toe flexor muscles were
consistently reduced during wedged walking. The redistribution of triceps surae loading from
the medial to lateral side has relevance in the rehabilitation of calf strains and tears, where
strains and tears are more common in the medial head of Gastrocnemius and where heel
wedges are also a prescribed treatment.40
The sensitivity of these secondary ankle plantar- flexor muscles to changes in inter-
segmental moments is of particular interest, both clinically and computationally. From a
computational perspective, the use of a static optimisation approach to derive muscle forces
inherently makes less efficient muscles more sensitive to changes in inter-segmental
moments compared to more efficient muscles. Tibialis Posterior and the toe flexors are
significantly less powerful and efficient ankle plantar- flexors compared to the triceps surae
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muscles, and as such these are the muscles that one would expect to respond to changes in
ankle moment. However, from a clinical perspective, it is not known if the body responds in
such a way to changes in demand at individual joints, possibly due to practical constraints of
determining deep muscle electromyography (EMG) data. If this observation is indeed found
to be true, it may provide some explanation as to why Achilles tendonitis patients respond so
variably to orthotics. Therefore, it is recommended that future research into the use of
orthotics to treat Achilles tendonitis focuses on the mechanical response specific orthotics
induce in the body at both individual muscle and whole joint levels with a view to proposing
methods to improve the conservative management of Achilles tendonitis.
A practical limitation of the study is that due to the inclined section being fixed to 2m,
it should be noted that some subjects participants may not have reached a steady-state of
inclined walking, despite using the middle of the three steps on the inclined surface.
However, the current setup does represent a change in walking incline which would be
commonly experienced in daily life. A second limitation of this study relates to the
lackabsence of direct measures or estimates of tendon strain, which would have
complimented the changes in muscle and tendon force observed. While estimating tendon
strain here would require combining a finite element tendon model with the musculoskeletal
model implemented here, which is beyond the scope of this study, such an approach may
have provided a greater depth of understanding into the mechanical changes that are induced
by orthotic heel wedges. relates to the lack of surface EMG data, specifically of the triceps
surae; as corroborating EMG data would have provide some measure of confidence in the
modelling approach used here. A final limitation of the study is related to the sensitivity of
musculoskeletal models to muscle morphology and geometry inputs. While a limitation of
any model that does not have MRI datasets of their subject participant cohort, the use of
scaled cadaveric muscle data in estimating muscle loads during various activities is a
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common approach,41-44 particularly when comparing intra-subject gait patterns such as in this
study.
This study characterised the effect of orthotic heel wedges on lower limb
biomechanics in the context of level, uphill and downhill walking, with the intention of trying
to quantify any mechanical differences that may arise due to heel wedges and help explain
the current mixed evidence surrounding heel wedges as a treatment for Achilles tendonitis.
Heel wedges were unable to significantly alter hindfoot kinematics, but did result in changes
in inter-segmental moments at all joints of the lower limb. However, heel wedges were
unable to reduce Achilles tendon loading across level, uphill and downhill walking conditions
in healthy individuals. Instead, secondary ankle plantarflexor muscles consistently showed
substantial reductions in loading. These results add to the body of largely clinical evidence
(pain, functional improvement etc.) that indicates that heel wedges are not an appropriate
treatment for Achilles tendonitis and has provided evidence as to why this is the case.
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Figure 1: Image of the 12mm orthotic heel wedge used by all subjectsparticipants
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Figure 2: Sketch of the inclined walkway setup
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Figure 3: Posterior (left) and lateral (right) views of the optical marker setup. Note: Grey circles indicate reflective marker positions
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Figure 4: Mean shod (black lines) and wedged (grey lines) joint ankle joint angles (A-C) and moments (D-F)kinematics when walking uphill (left column), on level ground (middle column) and downhill (right column) for the knee (A-C), ankle (D-F) and MTP (G-I) joints. Note: Solid lines represent flexion/extension angles and dashed lines represent inv/eversion ankle angles and moments (or ab/adduction for the knee). * denotes a statistically significant difference
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D
GD
H I
FE
*†
*
*†
*†
†
†
*†*
*
A B C*
†
*†
*
*
*
*
Figure 5: Mean shod (black lines) and wedged (grey lines) joint moments when walking uphill (left column), on level ground (middle column) and downhill (right column) for the knee (A-C), ankle (D-F) and MTP (G-I) joints. Note: Solid lines represent flexion/extension moments and dashed lines represent inv/eversion ankle moments (or ab/adduction for the knee).* and † denote statistically significant differences in magnitude and time to peak respectively.
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A
D
G
J K L
IH
E F
CB
**
*
Figure 65: Mean shod (solid lines) and wedged (dashed lines) muscle forces when walking uphill (left column), on level ground (middle column) and downhill (right column) for the Achilles (A-C), triceps surae (D-F), inv/evertor (G-I) and toe (J-L) muscles. Note: Abbreviations are: GastMed/GastLat – Medial/Lateral heads of the Gastrocnemius, SolMed/SolLat – Medial and lateral portions of Soleus, PeroB/L and PeroT – Peroneus Brevis/Longus and Tertius, TibAnt/TibPost, Tibialis Anterior and Tibialis Posterior, EDL/EHL – Extensor Digitorum/Hallucis Longus, FDL/FHL – Flexor Digitorum/Hallucis Longus. * denotes statistically significant change in peak force (only shown for the triceps surae muscles for clarity).
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Table 1: Comparison of forceplate and spatio-temporal gait characteristics across all inclines, mean (SD).
Uphill Level DownhillShod Wedged Shod Wedged Shod Wedged
Stance time [s] 0.72 (0.08)
0.73 (0.10)
0.69 (0.08)
0.70(0.07)
0.63 (0.09)
0.65* (0.08)
Velocity [m/s] 1.20 (0.24)
1.18 (0.24)
1.24 (0.20)
1.21(0.20)
1.20 (0.24)
1.17 (0.24)
First GRFPeak [BW]
1.08 (0.10)
1.07* (0.09)
1.12 (0.08)
1.12(0.08)
1.25 (0.13)
1.25 (0.14)
Time to 1st peak [% stance]
26(3)
28*(2)
26(4)
27(3)
25(4)
26*(4)
Second GRF Peak [BW]
1.17 (0.10)
1.18 (0.13)
1.13 (0.08)
1.12(0.09)
0.95 (0.11)
0.95 (0.11)
Time to 2nd peak [% stance]
79(3)
80*(2)
79(2) 80 (2) 79
(7)79(7)
ROFD[BWs-1]
6.22 (1.78)
5.83* (1.36)
7.45 (2.50)
6.78(1.48)
9.80(3.63)
9.17(3.81)
Most Posterior CoP
[normalised]
0.03 (0.07)
0.01 (0.09)
0.01 (0.07)
-0.01 (0.08)
0.03 (0.09)
0.01 (0.08)
Most Anterior CoP
[normalised]
0.68 (0.08)
0.65* (0.08)
0.66 (0.07)
0.63 (0.07)
0.66 (0.07)
0.62* (0.07)
CoP travel per step
[normalised]
0.64 (0.10)
0.63 (0.10)
0.65 (0.09)
0.64 (0.10)
0.63 (0.11)
0.61 (0.10)
Note: ROFD denotes the rate of force development and is derived from heel-strike to the first ground reaction force peak. * denotes P < .05 between shod and wedged conditions
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Table 2: Comparison of peak muscle and ankle joint reaction forces during shod and wedged walking across all inclines. Note: For full muscle names, see Figure 6 notes; Data presented as mean (SD).
Uphill Level DownhillPeak Force
[BW] Muscle Shod Wedged p-value Shod Wedged p-value Shod Wedged p-value
TricepsSurae
SolMed 1.30 (0.26) 1.18 (0.34) 0.047 1.16 (0.27) 1.03 (0.29) - 0.96 (0.23) 0.87 (0.23) 0.040
SolLat 0.77 (0.35) 0.88 (0.41) - 0.55 (0.33) 0.59 (0.34) - 0.53 (0.23) 0.60 (0.26) -
GastMed 0.72 (0.15) 0.67 (0.18) 0.040 0.66 (0.16) 0.59 (0.16) - 0.53 (0.12) 0.49 (0.12) -
GastLat 0.24 (0.08) 0.25 (0.08) - 0.20 (0.07) 0.20 (0.06) - 0.17 (0.04) 0.18 (0.05) -
Achilles 2.98 (0.69) 2.92 (0.81) - 2.47 (0.61) 2.33 (0.59) - 2.15 (0.48) 2.10 (0.54) -
Invertor/ Evertors
(1st half of stance)
PeroB/L 0.10 (0.11) 0.30 (0.30) 0.002 0.29 (0.27) 0.61 (0.45) 0.027 0.38 (0.30) 0.67 (0.51) 0.003
PeroT 0.02 (0.03) 0.07 (0.06) 0.002 0.09 (0.06) 0.15 (0.09) 0.035 0.09 (0.07) 0.15 (0.10) <0.001
TibAnt 0.26 (0.11) 0.38 (0.24) 0.058 0.53 (0.22) 0.79 (0.37) 0.030 0.56 (0.25) 0.75 (0.36) 0.002
TibPost 0.33 (0.15) 0.20 (0.14) <0.001 0.31 (0.13) 0.20 (0.16) 0.029 0.37 (0.21) 0.26 (0.22) 0.003
Invertor/ Evertors
(2nd half of stance
PeroB/L 0.09 (0.10) 0.17 (0.19) - 0.06 (0.10) 0.11 (0.13) - 0.06 (0.08) 0.10 (0.10) 0.083
PeroT 0.01 (0.00) 0.01 (0.01) 0.057 0.00 (0.01) 0.01 (0.01) - 0.01 (0.08) 0.01 (0.01) 0.076
TibAnt 0.31 (0.14) 0.21 (0.16) 0.012 0.33 (0.15) 0.26 (0.16) - 0.26 (0.10) 0.18 (0.10) 0.006
TibPost 0.50 (0.21) 0.36 (0.23) 0.020 0.53 (0.23) 0.41 (0.26) 0.095 0.41 (0.16) 0.29 (0.17) 0.005
Toes
EDL 0.02 (0.03) 0.06 (0.06) 0.002 0.09 (0.06) 0.15 (0.09) 0.036 0.09 (0.07) 0.15 (0.09) <0.001
EHL 0.02 (0.02) 0.06 (0.05) 0.002 0.08 (0.05) 0.14 (0.08) 0.032 0.09 (0.06) 0.14 (0.08) <0.001
FDL 0.04 (0.02) 0.03 (0.02) 0.023 0.04 (0.02) 0.03 (0.02) 0.058 0.04 (0.02) 0.03 (0.02) 0.003
FHL 0.39 (0.17) 0.29 (0.16) 0.015 0.42 (0.18) 0.31 (0.18) 0.064 0.34 (0.14) 0.25 (0.15) 0.002
Ankle JRF 5.17 (0.86) 4.88 (1.12) - 4.69 (0.83) 4.28 (0.85) - 3.82 (0.64) 3.53 (0.70) 0.037
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